Apparatus and methods for automatic optimization of interventricular and atrio-ventricular delays in real time for cardiac reynchronization in an active implantable medical device

ABSTRACT

An active implantable medical device for cardiac resynchronization with automatic and almost in real time optimization of the interventricular and atrio-ventricular delays is disclosed. The active implantable medical device includes a closed-loop for continuously controlling the atrio-ventricular delay AVD and the inter-ventricular delay VVD according to a hemodynamic signal delivered by a hemodynamic sensor. The closed-loop provides controlled modulation ( 38 ) and demodulation ( 42 ) the AVD, and modulation ( 48 ) and demodulation ( 52 ) the VVD, the modulation and demodulation being functionally interdependent ( 54, 56 ) by a sequence of alternating operation. A closed-loop regulator ( 36 ) for controlling the AVD receives as input an error signal (E AVD ) delivered based on demodulating the AVD ( 42 ) and outputs an AVD signal. A closed-loop regulator ( 46 ) for controlling the VVD receives as input a signal error (E VVD ) based on demodulating the DVV and outputs a VVD signal. In one embodiment, the regulators are PID controllers.

The present application claims the benefit of French Application No.10-50876 entitled “Active Implantable Medical Device for CardiacResynchronization with Automatic and Real Time Optimization of theInterventricular and Atrio-Ventricular Delays” and filed Feb. 9, 2010,which is hereby incorporated by reference in its entirety.

FIELD

The present invention relates to “active implantable medical devices” asdefined by the 20 Jun. 1990 Directive 90/385/EEC of the Council of theEuropean Communities, and more particularly to devices that continuouslymonitor a patient's heart rhythm and if necessary deliver to the heartof a patient electrical pulses for joint and permanent stimulation ofthe left and the right ventricles, so as to resynchronize them, saidtechnique being known Cardiac Resynchronization Therapy as (“CRT”) orBi-Ventricular Pacing (“BVP”).

BACKGROUND

Pacemakers providing CRT to a patient are known. One such device isdisclosed, for example, in EP 1108446 A1 and its counterpart U.S. Pat.No. 6,556,866 (assigned to Sorin CRM, previously known as ELA Medical),which is incorporated herein by reference and describes a deviceapplying, between the respective instants of stimulation to the left andthe right ventricles, a delay known as a variable interventricular delay(VVD). The VVD is adjusted so as to resynchronize the contraction ofboth left and right ventricles with fine optimization and improve thehemodynamic status of the patient.

A simultaneous stimulation of both left and right ventricles is notalways optimal because it does not necessarily lead to a synchronouscontraction of the ventricles. The delays of conduction in the left andright ventricular myocardium are not the same and may depend on multiplefactors such as the location of the ventricular leads and the type ofthe ventricular leads (e.g., an over-the-wire lead implanted into thecoronary sinus or an epicardial lead). It is therefore desirable toestablish a delay between the two stimuli (i.e., VVD) and to adjust thedelay VVD to resynchronize the contractions of the ventricles ensuring afine optimization of hemodynamic parameters. Depending on the patient'sclinical status, the VVD may be set to zero, positive (the leftventricle is stimulated after the right ventricle) or negative (theright ventricle is stimulated after the left ventricle).

A typical CRT pacemaker device runs a classic “dual chamber” operatingmode in which the device monitors the ventricular activity after anatrial event that is either spontaneous (P wave detection of an atrialdepolarization) or stimulated (application of an A pulse of atrialpacing). After detecting an atrial event, the device starts to count adelay period called “atrio-ventricular delay” (AVD) such that if nospontaneous ventricular activity (R wave detection of a ventriculardepolarization) is detected after this AVD, then the device triggers astimulation of the ventricle (application of a V pulse of ventricularpacing).

Clinical studies have demonstrated often dramatic improvements forpatients with congestive heart failure (“CHF”) whose condition was notimproved by conventional CHF therapies by precisely adjusting theparameters of the CRT therapy according to the patient's clinicalcondition and to the nature of specific myocardial contraction disorderssuch as dilated heart chambers, low ejection fraction, and excessivelengthening of the duration for the QRS complex (whether this disorderis spontaneous or induced by a traditional stimulation).

It is also necessary to reassess these parameter settings to optionallyre-adjust them if necessary: indeed, the benefits provided by the CRTtherapy may eventually lead to modifying the original configuration andthe parameter settings.

The standard technique for adjusting CRT stimulation parameters startswith the estimation of the characteristic delays of the systole byechocardiography, especially the timing of opening of the aortic valve.However, this procedure should be implemented in hospitals and byqualified personnel. This procedure is long and costly, thus cannot beapplied as often as it would be useful or necessary without interferingwith the patient's daily life, despite the beneficial effects on the CRTtherapy.

Another inherent difficulty with the echocardiographic assessment isthat it requires testing several successive pacing configurations, anddetermining an optimal AVD for each pacing configuration. Therefore, thenumber of combinations to be tested is important, and involves acomplicated and time consuming procedure that is difficult to manageexcluding it from a routine operation.

Therefore, there is a need for a technique for evaluating in a simple,rapid, automated and precise procedure the impact of different CRTstimulation parameters, including the AVD and VVD delays, so as tooptimize the patient's hemodynamic status.

One automated method of optimization is described in the article by J MDupuis, et al.: Programming Optimal Atrioventricular Delay in DualChamber Pacing Using Peak Endocardial Acceleration: Comparison with aStandard Echocardiographic Procedure, PACE 2003; 26: [Pt II], 210-213.This technique involves, while scanning the AVD in a given stimulationsetup to trace a characteristic, generating a value of the peak ofendocardial acceleration (“PEA”) according to the AVD. The consideredoptimal value of the AVD is the inflection point of the characteristics,i.e., the point corresponding to a maximum duration of ventricularfilling without truncating the A wave (i.e., a minimum delay between theclosing of the mitral valve and the beginning of the QRS complex).

Even if the corresponding algorithm gives satisfactory results, it takesup to several minutes before an optimal pair {AVD, VVD} is selectedbecause multiple scans of AVD are required for various values of VVDthat are separately tested.

Another drawback of this optimization technique is that the search foreach delay period (AVD or VVD) is independent of the other: for a givenVVD, a scan of the AVD gives a locally optimal AVD, but (as is explainedin greater detail below with particular reference to FIG. 5) theconvergence to a local optimum does not necessarily lead to the globaloptimum. In other words, the optimal pair {AVD, VVD} does notnecessarily correspond to an optimal AVD value, at constant VVD, or toan optimal VVD value, at constant AVD.

A special technique of optimization, which is faster, thus implementablein real time, is described in WO 2006/090397 A2 and WO 2006/126185 A2.The algorithm described in these documents use a spike neural network toidentify the maximum of a hemodynamic function (stroke volume). A spikenetwork, however, requires a dedicated processor, thus involving thedesign of a specific, more complex device demanding higher powerconsumption. A software implementation of this algorithm is possible,but in such a case, it requires computing resources that areunattainable in ultra-low power consumption microcontrollers adequatefor use in implantable medical devices.

WO 2008/010220 describes yet another technique, in which a spike neuralprocessor is combined with a reinforced learning algorithm (Q-learning),which learns and associates the cardiac conditions to the optimaldelays. Using this Q-learning can offer improved immunity to noise andincrease the speed of convergence of the control algorithm. However, inorder to achieve this performance, additional hardware resources arerequired including a microprocessor in addition to the spike neuralprocessor, which incurs extra cost, higher power consumption, andincreased spatial requirements for an implant device.

OBJECT AND SUMMARY

It is, therefore, an object of the present invention to provide a new,simple, rapid, automated and efficient technique for simultaneouslyoptimizing both AVD and VVD parameters, despite the interdependentnature of these two parameters.

It is further an object of the present invention to apply theoptimization technique almost in real time, preferably with a responsetime of only a few cardiac cycles, and implementing the optimizationtechnique in simple material and software resources that are availablein a current implantable device such as a CRT pacemaker.

According to one embodiment of the present invention, a parameter pair(couple) {AVD, VVD} is optimized exploiting a hemodynamic surface, i.e.,a function of two variables Z=f (AVD, VVD) rather than separatelyutilizing the two distinct characteristics (i.e., two distinct functionshaving each of the two variables such as Z₁=f₁(AVD) and Z₂=f₂(VVD)).This hemodynamic surface represents the combinational characteristics ofthe entire system including the CRT device, the patient's heart, and thehemodynamic sensor, thus the characteristics varies with the current AVDand VVD values programmed into the CRT device, the patient's condition,and the type of hemodynamic sensor used.

In one embodiment of the present invention, a digital closed-loop systemhaving a conventional proportional-integral-derivative (PID) digitalcontroller is used for monitoring or tracking an optimal point on thehemodynamic surface. The implementation of the digital closed-loopsystem in the CRT device requires only incremental material and/orsoftware resources without incurring extra cost and design for anadditional hardware.

The present invention is generally directed to an active medical device,such as an implantable device for cardiac resynchronization bybiventricular pacing, comprising (i) means for detecting atrial andventricular events; (ii) means for stimulating the right and leftventricles; (iii) a sensor delivering a signal representative of apatient's current hemodynamic parameter; (iv) means for delivering tothe stimulation means an atrio-ventricular delay AVD, calculated fromthe moment of detection of a spontaneous or paced atrial event and afterwhich a stimulation of the right ventricle is delivered in the absenceof a detected spontaneous ventricular event; (v) means for delivering tothe stimulation means an inter-ventricular delay VVD between respectivetimes of stimulation of right and left ventricles, and (vi) aclosed-loop controller, continuously monitoring the AVD and the VVDaccording to the hemodynamic signal delivered by the sensor.

In one embodiment of the present invention, the closed-loop controllerresponds to the signal delivered by said at least one sensor andgenerates an AVD error signal representative of a difference between thecurrent value of the AVD and an optimal value of AVD; and a VVD errorsignal representative of a difference between the current value of theVVD and an optimum value of VVD. The closed-loop controller furtherincludes an AVD closed-loop regulator, receiving as input said AVD errorsignal and delivering as output an AVD signal; and a DVV closed-loopregulator, receiving as input said VVD error signal and delivering asoutput a VVD signal.

In a preferred embodiment, the AVD and VVD closed-loop regulators arePID regulators.

Preferably, generating the AVD or VVD error signal is obtained by acontrolled modulation and demodulation of the AVD or VVD, respectively.

In one embodiment, the closed-loop controller has (i) means forgenerating the AVD error signal according to the signal delivered by theat least one sensor, (ii) means for generating the VVD error signalaccording to the signal delivered by the at least one sensor, (iii)means for modulating and demodulating the AVD, and (iv) means formodulating and demodulating the VVD. The two means of generation oferror signals for the AVD and the VVD, the respective means formodulating and demodulating the VVD and the AVD, are functionallyinterdependent. In this embodiment, it is particularly preferable toprovide that the means for modulating and demodulating in a controlledmanner the VVD and the AVD, as well as the regulators for controllingthe AVD and the DVV, operate alternatively in the control of the AVD andthe VVD respectively, for a predetermined number of cardiac cycles, suchthat the regulator for controlling the AVD is inoperative during themodulation/demodulation of the VVD, and vice versa.

BRIEF DESCRIPTION OF THE DRAWINGS

Further features, characteristics and advantages of the presentinvention will become apparent to a person of ordinary skill in the artfrom the following detailed description of preferred embodiments of thepresent invention, made with reference to the drawings annexed, in whichlike reference characters refer to like elements and in which:

FIG. 1 is a block diagram of a system for the closed-loop, real-timehemodynamic CRT device, according to one embodiment;

FIG. 2 is a representation of a hemodynamic surface, in the case of asensor measuring the differences between the systolic and the diastolicpressures;

FIG. 3 is a representation of a hemodynamic surface, in the case of asensor measuring the value of the peak of endocardial acceleration(PEA);

FIG. 4 illustrates a PID loop controller applied to a closed-loophemodynamic of a CRT device;

FIG. 5 illustrates the case of a separate optimization of the AVD and ofthe VVD independently, without interaction;

FIG. 6 shows a variation in the hemodynamic function according to theAVD or to the VVD, in the case of a sensor measuring the differencesbetween the systolic and the diastolic pressures;

FIG. 7 illustrates a variation in the hemodynamic function according tothe AVD, in the case of a sensor measuring the value of the peak ofendocardial acceleration (PEA);

FIG. 8 shows a flow chart to obtain an optimization of the hemodynamicparameters using two interdependent PID loops withmodulation/demodulation of the AVD and of the VVD;

FIG. 9 is a timing diagram showing how the AVD and VVD parameters changeduring a search for an optimum, over successive cardiac cycles;

FIG. 10 illustrates a response of the closed-loop PID at each adaptationstep, for various gain settings of the loop;

FIG. 11 illustrates different possible characteristics of the sensorsignal as a function of the delay (AVD or VVD), to obtain an errorsignal for controlling the closed-loop;

FIG. 12 shows a flow chart to obtain an optimization of the hemodynamicparameter using two PID loops directly controlled by a sensor to obtaina direct measure of the error; and

FIG. 13 shows a flow chart to obtain an optimization of the hemodynamicparameter using two PID loops, one of which is controlled by a sensor todirectly obtain the error signal.

DETAILED DESCRIPTION OF THE INVENTION

With reference to FIGS. 1-13, various embodiments of the presentinvention will now be described. As regards its software aspects, thepresent invention can be implemented by an appropriate programming ofthe controlling software of a known device, for example, a cardiacpacemaker or a defibrillator/cardioverter, including means forcollecting signals provided by endocardial leads and/or one or moreimplanted sensors.

The present invention is directed to apparatus and method forimplementing the functions of automatic and in almost real timeoptimization of the AVD and VVD. One such device includes programmablemicrocontroller and/or microprocessor circuitry to receive, format,process electrical signals collected (detected) by one or more implantedelectrodes, and deliver stimulation pulses to these electrodes. It ispossible to transmit by telemetry software and store it in a memory ofthe implantable device to execute the functions of the present inventionthat will be described herein. The adaptation of these devices toimplement the functions and features of the present invention isbelieved to be within the abilities of a person of ordinary skill in theart, and therefore will not be described in detail. One suitable deviceto which the present invention may particularly be applied are those ofthe Paradym CRT device family, produced and marketed by Sorin CRM,Clamart France, formerly known as ELA Medical, Montrouge, France.

The various elements involved in the present technique of closed-loopreal-time hemodynamic optimization of the AVD and VVD parameters inaccordance with the invention are illustrated in a schematic functionalblock in FIG. 1. The reference 10 denotes the generator of the CRTdevice connected to the heart via leads to collect depolarizationsignals of the myocardium and stimulate the myocardium by deliveringelectrical pulses to the different cavities of the heart. A lead 12implanted in the right atrium (RA), and a lead 14 implanted in the rightventricle (RV) allows optimizing the atrio-ventricular delay AVD betweenthe instants of stimulation of the right atrium and the right ventricle.A lead 16 implanted inside or in the vicinity of the left ventricle(LV), in combination with the lead 14 implanted in the right ventricle,allows optimizing the interventricular delay VVD between the left andright ventricles.

A hemodynamic sensor 18 measures hemodynamic signals representingcardiac output from the heart. More specifically, the hemodynamic sensor18 estimates the changes in contractility, correlated with increases inblood pressure. Hemodynamic sensors differ from activity sensors (forexample, acceleration sensors) or metabolic sensors (for example, minuteventilation sensors) that are intended to diagnose the presence or levelof an activity by the patient and to quantify the patient's metabolicneeds. Depending on the patient's level of activity or metabolic needs,the stimulation heart rate is adapted. However, the hemodynamic sensor18 not only monitors the patient's efforts, similarly to metabolic oractivity sensors, but also gives an indication of the patient'shemodynamic tolerance in certain events, especially the tolerance to achange in the AVD and VVD parameters by the device.

In a preferred embodiment, the hemodynamic sensor 18 is an intracardiacpressure sensor that measures the pressure difference Δ_(pp) between thesystolic and the diastolic pressures, or an endocardial accelerationsensor that is capable of detecting Peak Endocardial Acceleration (PEA)signals.

These examples of hemodynamic sensors are, however, in no waylimitative, and the present invention may be implemented with othertypes of hemodynamic sensors such as: an epicardial acceleration sensor(not endocardial), a cardiac wall motion sensor, an intracardiacbioimpedance sensor, an optical oxygen saturation sensor, and anultrasound sensor for measuring changes in blood volume.

It should be understood that the hemodynamic signals used in theanalysis for the optimization of the AVD and VVD may be obtained from anexternal sensor instead of an implanted sensor, for example, through anaccelerometer sensor attached to the patient's chest at the sternum.

For various descriptions of hemodynamic sensors, reference may beparticularly made to the following documents, which are incorporatedherein by reference:

-   -   an endocardial acceleration sensor of the PEA type: EP 0515319        A1 and its counterpart U.S. Pat. No. 5,304,208, assigned to        Sorin Biomedica Cardio SpA describe a method to collect        endocardial acceleration signals using a lead provided with a        distal endocardial stimulation electrode implanted at the apex        of the ventricle and measuring the endocardial acceleration        using a micro-accelerometer. EP 0655260 A1 and its counterpart        U.S. Pat. No. 5,693,075, also assigned to Sorin Biomedica Cardio        SpA describe a method to treat the measured endocardial        acceleration signals to notably derive endocardial acceleration        peak values corresponding to the two major noises that are        identifiable in each heart cycle of a healthy heart;    -   a transvalvular bioimpedance sensor measured between the atrium        and the ventricle located on the same side of the heart: EP        1116497 A1 and its counterpart U.S. Pat. No. 6,604,002 assigned        to Sorin CRM S.A.S previously known as ELA Medical describe a        dynamic measurement of bioimpedance (BioZ) to assess the        diastolic and systolic volumes and obtain an indication of the        cardiac output and thus, an indication of the ejection fraction;    -   a transseptal bioimpedance sensor measured between a site        located on one side of the heart and a site located on the other        side: EP 1138346 A1 and its counterpart U.S. Pat. No. 6,725,091        assigned to Sorin CRM S.A.S previously known as ELA Medical        describe another type of measure useful to deliver a        representative value of the ejection fraction; and    -   an external accelerometer sensor: EP 1741387 A1 and its        counterpart U.S. Pat. No. 7,613,507 assigned to Sorin CRM S.A.S        previously known as ELA Medical describe the method to collect        endocardial acceleration signals using a lead placed on the        patient's chest.

Whether the hemodynamic sensor is implanted (transvenous, epicardial . .. ) or externally placed, it delivers signals correlated with thecardiac output and transmits the signals to an acquisition circuit 20.Acquisition circuit 20 preferably is incorporated in the generator 10 ofthe device, or but alternatively, it may be separately packaged from thegenerator 10 and located outside the patient's body.

The acquisition circuit 20 delivers a signal Z, hereinafter referred toas a “signal” or a “hemodynamic signal”, to a closed-loop controller 22and more preferably a PID controller. The transmission from theacquisition circuit 20 to the controller 22 can be direct (in the caseof a treatment which is purely internal to the implant), or by telemetry(in the case of an external sensor and a controller 22 incorporated intothe implant, or in the case of an implanted sensor and of an externalcontroller 22 that is integrated into a programmer used for setting upthe generator during a visit to a practitioner).

The controller 22 implements a closed-loop algorithm to concurrentlyobtain optimal values of the AVD and of the VVD.

It should be understood that it is possible to use multiple sensors forcontrolling a delay (e.g., an accelerometer and a bioimpedance sensorfor controlling the AVD), or to use a different sensor for each delay(e.g., a bioimpedance sensor to control the VVD and a PEA sensor tocontrol the AVD or vice versa). In any event, whatever the type ornumber of sensors used, the acquisition circuit 20 delivers a piece ofinformation Z as a function of two variables Z=f (AVD, VVD). Thefunction Z can be graphically represented by a surface and hereinafterreferred to as a “hemodynamic surface”.

FIG. 2 illustrates an example of a hemodynamic surface Z=f (AVD, VVD).In this case, the hemodynamic sensor 18 is blood pressure sensordelivering signals to obtain a piece of information Z representative ofa difference Δ_(pp) between the systolic pressure and the diastolicpressure. The optimum of a hemodynamic surface 24 is located at thepoint Z_(opt) having the highest Z score. In practice, the vicinity ofthis optimum Z_(opt) in the hemodynamic surface 24 may be approximatedby parabolic characteristics in the plane {VVD, Z} and in the plane{AVD, Z}.

FIG. 3 illustrates another example of hemodynamic surface Z=f (AVD,VVD). In this case, the hemodynamic sensor 18 is an endocardialacceleration sensor delivering signals to obtain a piece of informationZ representative of the peak of endocardial acceleration PEA. Morespecifically, the first peak of endocardial acceleration (“PEA1”)corresponds to the closure of the mitral and tricuspid valves, at thebeginning of the phase of isovolumetric ventricular contraction(systole), and its variations are closely linked to the variations ofblood pressure in the ventricle representing the myocardialcontractility. The amplitude of the PEA1 peak is particularly correlatedto the positive maximum of the pressure variation dP/dt in the leftventricle.

The signal variations as a function of AVD follow a relation Z=f (AVD),represented by a characteristic curve 26 in the form of a sigmoid, whilethe variations of the same signal as a function of VVD follow a relationZ=f (VVD) represented by an approximately parabolic characteristic curve28. In this case, the optimal value of the couple {AVD, VVD} correspondsto the intersection point Z_(opt) between the inflection point of thesigmoid characteristic curve 26 and the highest point of the paraboliccharacteristic curve 28.

FIG. 4 illustrates an exemplary functional diagram of a closed-loop ofthe system of FIG. 1. A controller 22 outputs the two values AVD andVVD, from an input E_(Z) representing the error between the currentvalue Z=f (AVD, VVD) obtained from the value measured by the hemodynamicsensor 18 and a reference value corresponding to the desired optimumZ_(opt) of the hemodynamic surface 30. The hemodynamic surface 30characterizes the response Z=f (AVD, VVD) of the group consisting of:the CRT resynchronization generator 10, the patient's heart, thehemodynamic sensor 18 and the acquisition circuit 20.

In one embodiment, the controller 22 is a digital PID controller forclosed-loop control delivering any combination of the following outputcontrol signal components:

-   -   proportional component (“P”): the output is proportional to the        error signal E_(Z);    -   integral component (“I”): the output is the sum of the        instantaneous error signal E_(Z) over time. The output is        increased as the error E_(Z) accumulates over time, thus        eliminating the residual stead-state error, and    -   derivative component (“D”): the output is proportional to the        time derivative of the error E_(Z) creating an accelerated        response in case the error rate (dE_(z)/dt) increases.

In the case of a digital controller operating on discrete values, such aclosed-loop system works intermittently in successive steps (each stepin this case corresponding to a cardiac cycle). In this case, it isnecessary to wait to the end of the cardiac cycle k to obtain the valueZ_(k) that represents the value of the hemodynamic function during thecorresponding cardiac cycle.

A conventional PID controller may not be directly used to implement thepresent invention because:

-   -   the target point value Z_(opt) is not known or not available a        priori and    -   a conventional PID controller has a single input and a single        output (SISO), whereas in the present case it is necessary to        simultaneously monitor two variables, namely the AVD and the        VVD.

Nor is it possible to split the system so as to separately andsimultaneously control the AVD and the VVD with two PID controllersbecause duplicate controllers do not ensure convergence to an optimalpoint.

These drawbacks of using a conventional PID controller for theimplementation of the present invention are more specifically explainedwith reference to FIG. 5. The hemodynamic surface of FIG. 5 correspondsto the parabolic surface as shown in FIG. 2 where the parameter Z is afunction of the difference Δ_(pp) between the systolic pressure and thediastolic pressure. In a plane {AVD, VVD}, the contours correspondapproximately to ellipses, shown in dotted lines with a global optimumZ_(opt) at point P.

With a fixed value of AVD at AVD1, a scan of the VVD is performed alongthe axis S1. The maximum VVD is found at VVD1. If from this localmaximum VVD1, a scan of the AVD is performed along the axis S2, and alocal maximum point is found at point P1.

If, however, the value of the VVD is fixed at VVD2 and that the AVD isscanned along axis S3, a local maximum AVD is found at AVD2. From thislocal maximum AVD2, a scan of the VVD is performed along axis S4, andanother local maximum point is found at point P2.

Thus, the optimization of the VVD, then of the AVD, leads to an optimumat P1, while the optimization of the AVD and then the VVD leads toanother optimum P2. These two points P1 and P2 are not only differentfrom one the other, but also different from the real optimum located atP.

Therefore, an algorithm using two independent closed-loop systems, onefor the AVD and the other for the VVD, would not always find the globaloptimum.

A preferred process in accordance with the present invention, as will bedescribed below, comprises:

-   -   first, generation of an error signal to guide the digital PID        controller to the optimal point Z_(opt) and    -   second, use of the error signal thus generated to simultaneously        perform closed-loop assessments of the AV and VV delays.

Generation of the Error Signal

The method for generating the error signal is described with referenceto FIGS. 6 and 7. To determine the position of the current AVD and VVDvalues (i.e., values currently programmed into the device) compared tooptimal values AVD_(opt) and VVD_(opt) that correspond to the globaloptimum Z_(opt), the following steps are performed:

-   -   modulating the delays AVD and VVD, so that the delays vary in a        deterministic and known manner around their current value(s),        and    -   observing the changes resulting from the hemodynamic signal Z to        derive an error signal and extracting information about the        error signal by means of demodulation.

The modulation and demodulation apply separately to the AVD and VVD, orsimultaneously on both. To ensure patient comfort, however, anindependent modulation of the AVD and VVD is preferable. The choice ofmodulation and of the demodulation depends on the characteristics of thehemodynamic surface.

As shown in FIG. 6, in the case of a blood pressure sensor, thedifference Δpp between the systolic pressure and the diastolic pressureis illustrated as the hemodynamic parameter Z. Near the optimum Z_(opt),the hemodynamic surface Z is approximately a parabolic surface, both fora variation of the AVD and for a variation of VVD. The current value ofeither the AVD delay or the VVD delay is noted as X.

It is assumed that the function Z=H (X-X_(opt)) is an even functionaround X_(opt) that corresponds to the optimal delay value (either AVDor VVD). Specifically, any function with a similar maximum can beapproximated by a symmetric quadratic function in a vicinity of X_(opt)and, in any event, if the function is not perfectly symmetrical, thecontrol algorithm converges to a limit X′_(opt) close to X_(opt), thedifference having no significant impact on the final outcome from theclinical perspective for the patient.

The modulation of the delay X is performed by varying the delay X aroundits current value, with a modulation step of Δ milliseconds. The deviceis sequentially programmed with three different values of X (AVD orVVD):

-   -   X, the current value of the AVD (or VVD), which generates a        hemodynamic value Z₁=H (X-X_(opt));    -   X+Δ, which generates a hemodynamic value Z₂=H (X+Δ−X_(opt)); and    -   X−Δ, which generates a hemodynamic value Z₃=H (X−Δ−X_(opt)).

According to one embodiment, the demodulation is performed bymultiplying the measured signal by +1 or −1, depending on the directionof the modulation. In other words, the samples Z₂ and Z₃ are passed in afirst-order differentiator filter with two coefficients, +1 and −1,i.e., Z₂ is multiplied by +1 and Z₃ by −1. The error signal obtained atthe demodulator output is thus E=Z₂−Z₃.

In the case of an even function H, the error signal E can be written asE=2p (X_(opt)−X), p=−H′ (Δ)=H′ (−Δ), p being the slope of thehemodynamic curve at X=X_(opt)−Δ (or X=X_(opt)+Δ). Thus, the errorsignal obtained after modulation and demodulation is proportional to thedifference between the target delay X_(opt) and the current delay X. Thesensitivity of the error signal depends on the slope p of thehemodynamic curve Z. The slope of this curve is equal to zero whenX=X_(opt). When the modulation step Δ is small, the slope p is alsosmall. The selection of the size of the step Δ depends on the quality ofthe hemodynamic signal.

When the hemodynamic function is almost flat, the slope p is small, andthe error signal E is not very sensitive to variations in X. Thisscenario can be detected by estimating the curvature of the hemodynamicsurface using the measure Z₁. The curvature C is estimated by:C=2Z₁−Z₂−Z₃. When the curvature C is smaller (in absolute value) than apredefined threshold, it is preferable not to use the error signal E. Inthis case, the delay X is not adjusted (neither modulated nordemodulated), and the current delay X is used as the optimal delayX_(opt).

Some CRT devices allow programming of the delays with only a limitednumber of predefined values. With limited choices of modulation steps Δ,the modulation may not be symmetrical around the current delay X. Forthe modulation step Δ₂ associated with the measurement Z₂ and themodulation step ΔΔ₃ associated with the measurement Z₃, the error signalE′ is obtained by:

$E^{\prime} = {2{\frac{{\left( {\Delta_{2} - \Delta_{3}} \right)Z_{1}} + {\Delta_{3}Z_{2}} - {\Delta_{2}Z_{3}}}{\left( {\Delta_{2} + \Delta_{3}} \right)}.}}$

The error signal E′ is zero when X coincides with X_(opt).

As shown in FIG. 7, in case of an endocardial acceleration sensor, thehemodynamic parameter Z corresponds to the amplitude of PEA based on thevariation of the AVD. The hemodynamic parameter Z is of a sigmoid shape,and the optimum Z_(opt) does not correspond to an extremum (as in thecase of FIG. 6) but rather to a point of inflection of the sigmoid.

In this case, the method for generating the error signal is modified asfollows. It is assumed that near the optimum Z_(opt), the function Z canbe expressed as Z=Z_(opt)+G(X−X_(opt)), G being an odd function aroundX_(opt) and X being the current value of the AVD.

In general, a function with an inflection point can be approximated byan odd function G in a vicinity of the inflection point, in this caseX_(opt). If the function G is not completely odd, the control algorithmconverges to a delay X′_(opt) close to X_(opt), and the difference haveno significant impact on the final outcome from the clinical perspectivefor the patient.

The modulation is performed by varying X around its current value with amodulation step of Δ milliseconds. The device is sequentially programmedwith three different values of AVD:

-   -   X, the current value of the AVD, which generates an hemodynamic        value Z₁=Z_(opt)+G(X−X_(opt));    -   X+Δ, which generates an hemodynamic value Z₂=Z_(opt)+G        (X+Δ−X_(opt))    -   X−Δ, which generates an hemodynamic value Z₃=Z_(opt)+G        (X−Δ−X_(opt)).

The demodulation is performed by passing the samples Z₃, Z₁ and Z₂ in asecond-order differentiator filter with three coefficients, respectively−1, +2 and −1. In other words, Z₁ is multiplied by 2, Z₂ is multipliedby −1 and Z₃ is multiplied by −1.

In this case, the error signal obtained at the output of the demodulatoris written as: E=2Z₁−Z₂−Z₃. This error signal can be rewritten as:E=2(p₀−p₁)(X_(opt)−X), where the values p₀ and p₁ are those of the(positive) slopes p₁=−G′(Δ) and p₀=−G′(0).

Thus, the error signal obtained after modulation and demodulation isproportional to the difference between the target delay X_(opt) and thedelay to be controlled X.

When the hemodynamic value Z linearly varies with X, any point X gives azero error signal because the second derivative (which defines theinflection point) is zero everywhere when p₀=p₁.

The sensitivity of the error signal in the case of seeking an inflectionpoint thus depends on the difference p₀−p₁. Therefore, it is better tochoose a modulation step Δ large enough to maximize this sensitivity.When the hemodynamic function is almost flat or linear, the differencep₀−−p₁ is small, and the error signal E is insensitive to variations inX. In this case, it is preferable not to use the error signal E. In thiscase, the delay X is not adjusted (neither modulated nor demodulated),and the current delay X is used as the optimal delay X_(opt).

Some CRT devices allow programming of the delays (AVD or VVD) with alimited number of predefined values. With the limited choice ofmodulation steps Δ, the modulation may not be perfectly symmetricalaround the current delay X. If Δ₂ is the modulation step associated withthe measurement Z₂ and Δ₃ the modulation step associated with themeasurement Z₃, the error signal is obtained by:

$E^{\prime} = {2{\frac{{\left( {\Delta_{2} + \Delta_{3}} \right)Z_{1}} - {\Delta_{3}Z_{2}} - {\Delta_{2}Z_{3}}}{\left( {\Delta_{2} + \Delta_{3}} \right)}.}}$

The error signal E′ is equal to zero when X coincides with X_(opt).

The sequence of modulation/demodulation during a cardiac cycle isoperated as follows (either for the AVD or the VVD):

-   -   a) measuring hemodynamic signal Z₁ for a value X of the AVD or        the VVD,    -   b) modifying the delay: X=X+Δ,    -   c) measuring the resulting hemodynamic signal Z₂,    -   d) modifying the delay: X=X−Δ,    -   e) measuring the resulting hemodynamic signal Z₃,    -   f) calculating the error signal E=Z₂−Z₃ (or E=2E₁−Z₂−Z₃, as        appropriate).

In one embodiment, the signals Z₁, Z₂ and Z₃ at steps a), c) and e) aremeasured either over one cardiac cycle, or by taking an average overseveral cardiac cycles depending on the response time of the measuredparameter to the changes in the AVD or VVD. For some hemodynamicsensors, it may be necessary to wait for one or more cardiac cycles sothat the hemodynamic response is stabilized after modifying the delay insteps b) and d).

Simultaneous Closed-Loop of the AVD and of the VVD from the Error Signal

With reference to FIGS. 8-10, a preferred algorithm in accordance withthe present invention for using the AVD and VVD error signals generatedby the by the modulation/demodulation described above, will now bedescribed.

In one embodiment, the classical theory of closed-loop control systemsis used to realize a simple digital controller to monitor the optimalAVD and VVD.

As shown in FIG. 8, the system includes two closed-loop systems for eachof the two delays AVD and VVD. Each of these closed-loops includes arespective dedicated controller 36, 46, including a PID controller(e.g., a non-restrictive type) associated with the error generationblock 34.

The closed-loop of the AVD includes a PID controller 36 receiving asinput an error signal E_(z) related to the AVD and delivering as outputa control signal applied to a modulator circuit 38 for, as explainedabove with reference to FIG. 6, modulating in a controlled manner thevalue of AVD around the optimal value to be sought.

Block 40 represents the system responsive to a variation of the AVD. Theoverall system includes (see FIG. 1) the CRT generator 10, the patient'sheart, the hemodynamic sensor 18, and the acquisition circuit 20. Thisblock 40 incorporating the function represented by the hemodynamicsurface 30 generates hemodynamic signal Z. The hemodynamic signal Z isdemodulated by block 42, and the error signal E_(z) is delivered as aninput to the PID controller 36.

Similarly, the closed-loop of the VVD includes a PID controller 46receiving as input an error signal E_(z) related to the VVD anddelivering as output a signal of control of the VVD applied to amodulator circuit 48 for, as explained above with reference to FIGS. 6and 7, modulating in a controlled manner the value of the VVD around theoptimal value to be sought. The block 40 provides hemodynamic signal Z.The hemodynamic signal is demodulated by the block 52, and the errorsignal E_(z) is delivered as an input to the PID controller 46.

As explained above with reference to FIG. 5, an independent research ofthe optima for the AVD and the VVD generally leads to local optima thatare different from the global optimum to be sought.

The two interdependent closed-loops interact between the blocks ofmodulation and of demodulation of the AVD and VVD, as shown by arrows 54and 56. The interaction may in particular result from a specificsequence of the cycles of modulation/demodulation of the AVD and theVVD, respectively.

An example of a specific sequence is shown in FIG. 9. As noted abovewith respect to the generation of the error signal, the error signal Eis available after demodulating the values Z₁, Z₂ and Z₃ that correspondto the successive respective values X, X+Δ and X−Δ of the delay X (AVDor VVD). Therefore, at least three cardiac cycles are required to obtainan error signal for each of the AVD or VVD delays.

The PID controller 36 (or 46) sequentially operates an update of thedelay after multiplication of this error signal by a gain, and thenintegration. For the AVD, the corresponding PID controller 36 waits forthree cardiac cycles to receive an error signal. It then performs anupdate of the AVD controlled at its output, and remains idle for thefollowing three cardiac cycles, during which the PID controller 46 forthe VVD takes over, and so on, alternating the operation of the two PIDcontrollers 36 and 46 ensuring the convergence and the simultaneousmonitoring of both optimal delays, AVD_(opt) and VVD_(opt).

In FIG. 9, the values of the AVD and VVD delays during a sequence ofeight cardiac cycles, are illustrated. During cycles #1 to #3, the AVDis modulated and the VVD is maintained at its current value (inhibitionof modulation of the VVD). At the end of cycle #3, an error signal C1 iscalculated for the AVD. This signal C1 is applied at the input of thePID controller 36 dedicated to the AVD, and a new value of the AVD is sodetermined as an output of the controller 36.

The VVD is then modulated during the cycles #4 to #6, keeping unchangedthe value of the AVD just obtained (that is to say that the modulationof the AVD is inhibited during these three cycles). At the end of cycle# 6, an error signal C2 of the VVD is calculated and applied to the PIDcontroller 46 dedicated to the VVD, providing as output a new value ofthe VVD applied to the device (Cycle #7). In cycle #8, the describedprocess is repeated by applying an AVD modulation, and so on. The cycles#7, #8 and #9 correspond to the cycles #1, #2 and #3 of the succeedingiteration. A complete cycle of adapting the system to the new values ofthe pair AVD/VVD therefore requires at least six cardiac cycles.

The condition for stability and convergence of the dual closed-loopsystem as described above is evaluated as follows. Let k be the timeindex for a PID controller (the controller 36 for the AVD or thecontroller 46 for the VVD). A cycle of adaptation of the closed-loopsystem corresponds to the passage from k to k+1 that is a duration of atleast six cardiac cycles.

The error signal provided to the PID controller is written:

E _(k) =s(X _(opt) −X _(k)),

where s is the sensitivity of each controller that depends on themodulation step Δ and the shape of the hemodynamic function.

The PID controller corrects X_(k) according to the equation:

X _(k+1) =X _(k) +gE _(k),

where g is the gain of the loop.

The z transform of the transfer function K between the currentcontrolled delay X and the optimum delay X_(opt) is expressed as:

$\frac{X}{X_{opt}} = {{K(z)} = \frac{\beta}{z - 1 + \beta}}$

in which β=g.s.

The behavior of the closed-loop in the long term is determined by thelimit of this function when z tends to 1. As lim K (z)=1 and as the loopis closed, the controlled delay X converges to X_(opt) provided that thestability condition of the closed-loop is verified. This condition canbe expressed as:

0<β<2

0<g<2/s

The loop gain thus depends on the sensitivity s of the controller.

FIG. 10 shows the evolution of the controlled delay X in response to astep in X_(opt) of 10 ms for three values of β: for β=1, the controlleddelay X coincides with the optimum time after a single cycle ofadaptation, while for β<1 the system monotonically converges and for β>1the controlled delay hovers around X_(opt).

In practice, the value of the sensitivity s is not known. It can beestimated for a given patient and a given sensor, at least once, forexample, at the time of implantation of the sensor. In any event, it isnevertheless preferable to choose a gain smaller than 1/s, to avoidoscillations and allow a margin of stability if the sensitivity sincreases over time.

In another embodiment, stability is ensured by applying a closed-loopalgorithm of the “PID truncated” type with limited adjustments to agiven maximum value c_(k)=min (20, max (−20, gE_(k)), for example,limiting adjustments to ±20 ms.

A closed-loop algorithm of the “PID truncated” type is applied to thecase of an hemodynamic signal delivered by the PEA sensor in thefollowing steps:

-   -   a) Perform a complete scan when the sensor is installed    -   b) Estimate the slope p for the VV curve, p₀, and p₁ for the AV        curve, the minimum slope P_(threshold) for AV and the minimum        curvature C_(r) threshold for VV    -   c) Select the gain g₁=0.35/p for the VV loop    -   d) Select the gain g₂=0.35/(p₀−p₁) for the AV loop    -   e) AV_(p)=120 ms, VV_(p)=0 ms, Δ₁=10 ms, Δ₂=15 ms    -   f) Program AV=AV_(p) and VV=VV_(p)    -   g) Start the servo loop:

1. Measure Z₁

2. Program AV=AV_(p)+Δ₂

3. Measure Z₂

4. Program AV=AV_(p)−Δ₂

5. Measure Z₃

6. E=g2*(2*Z₁−Z₂−Z₃)

7. P=Z₃−Z₂

8. C=min(20, max(−20, E))

9. If P>P_(threshold)

10. Program AV=AV_(p)

In an alternative implementation of the algorithm, a control algorithmof the “PID sign” type, with a fixed step, of 5 ms, for example, isapplied. In this case, each PID controller performs an update accordingto the relationship X_(k+1)=X_(k)+5*sign (E_(k)), in which sign ( . . .) represents the sign function (equal to 1 if E_(k) is positive and to−1 otherwise). Stages 6-9 of the above algorithm, then come to:

1. E=2*Z₁−Z₂−Z₃

2. AV_(p)=AV_(p)+sign(E)*5

This variant is advantageously robust, however has a disadvantage ofperpetually oscillating around the optimum time.

Variants of Implementation of the Invention

The implementation of the present invention is not limited to caseswhere the error signal is obtained by a technique ofmodulation/demodulation. Specifically, FIG. 11 illustrates several typesof signal Z delivered by the hemodynamic sensor 18 depending on thevariation of the X delay (AVD or VVD). FIG. 11 (a) corresponds to thecase illustrated in FIG. 6, in which the optimum Z_(opt) is located atthe curve (supposed to be a pair at this point) maximum, and the optimumis possibly reached by successive approximations using the technique ofmodulation/demodulation as described above. This characteristic curve isobtained from the difference Δ_(pp) between the systolic pressure andthe diastolic pressure that are measured by a blood pressure sensor.

FIG. 11 (b) corresponds to a characteristic of the same type, in which aminimum of the curve is sought. The minimum is supposed to be pair inthe vicinity of the optimum Z_(opt). This characteristic curve istypically obtained by measuring the width of the QRS complex on the ECG.It may be referred in particular to David Tamborero and al.,Optimization Of The Delay in Cardiac Resynchronization interventricularTherapy Using The QRS Width, American Journal of Cardiology, 15 Nov.2009, Vol. 104, Issue 10, pp. 1407-1412, which describes an optimizationprocess in which the optimum VVD delay corresponds to the narrower QRScomplex.

FIG. 11 (c) corresponds to the case illustrated in FIG. 7, with acharacteristic of the sigmoid type, which is typical for a signaldelivered by an endocardial acceleration sensor, such as a sensor givingthe value of the PEA peak based on the AVD. In this case, the optimumZ_(opt) is at the inflection point of the curve, assumed to be odd nearthe inflection point, and the error signal is generated by a techniqueof modulation/demodulation as described above.

In the case of FIG. 11 (d), however, the sensor delivers a signaldirectly proportional to the error signal, in which the optimum Z_(opt)is located at the origin, and the signal Z=k*E_(x) is substantiallyproportional to the sought error signal E_(x). It is not necessary inthis case to operate by modulation/demodulation, insofar as the errorsignal is directly readable at the output of the sensor. Thischaracteristic curve is obtained, for example, from a sensor of the“zero crossing” type such as a differential bio-impedance sensordelivering a signal depending on the phase difference between thebioimpedance measured in the left ventricle (LVZ) and the one measuredin the right ventricle (RVZ). U.S. 2008/0114410 describes, among otherthings, a system for optimizing the AV delay of a CRT device bymeasuring the differential bioimpedance.

FIG. 12 is a counterpart of FIG. 2, in a scenario in which both of theerror signals E_(AVD) and E_(VVD) of the AVD and the VVD, respectively,can be obtained directly from the signal Z delivered by the hemodynamicsensor 18. In this case, the signal Z is applied to the input of the PIDcontroller 36 for controlling the AVD, and in the same manner, to theinput of PID controller 46 for controlling the VVD.

FIG. 13 illustrates an intermediate case in which the error signal ofone of the delays, for example, the error signal E_(AVD) of the VVD, isdirectly obtained from the signal Z delivered by the hemodynamic sensor18, while the signal error E_(VVD) of the VVD to be applied to the inputof the PID controller 46 for control of the VVD requires amodulation/demodulation of the VVD.

In both cases of FIGS. 12 and 13, it is necessary to have two separatePID controllers for both AVD and VVD delays. However, insofar as it ispossible to directly obtain an error signal for at least one of thedelays, there is no need to alternate periods of modulation/demodulationon each of two loops, as it had been described with reference to FIGS. 9and 10, with alternating periods of inactivity of one of the PIDcontrollers while generating the error signal for the other loop.

One skilled in the art will appreciate that the present invention may bepracticed by other than the embodiments described above, which arepresented for purposes of illustration and not of limitation.

1. An active implantable medical device for resynchronization bybiventricular pacing, comprising: means for detecting atrial andventricular events; means for stimulating right and left ventricles; atleast one sensor (18) delivering a hemodynamic signal (Z) representativeof a patient's current hemodynamic parameter; means for applying to themeans for stimulating an atrio-ventricular delay AVD, the AVD beingcounted from the detection of a spontaneous or paced atrial event untila pacing of the right ventricle is applied in the absence of a sensedspontaneous ventricular event; means for applying to the means forstimulating an inter-ventricular delay VVD between the respectivemoments of stimulation of the right and left ventricles; and closed-loopmeans (22), for continuously monitoring the AVD and the VVD according tothe hemodynamic signal delivered by the at least one sensor, means forgenerating an error signal of AVD (34) according to the hemodynamicsignal delivered by said at least one sensor (18), the error signal ofADV (E_(AVD)) being representative of a difference between the AVD andan optimum value of the AVD (AVD_(opt)); means for generating an errorsignal of VVD (34) according to the hemodynamic signal delivered by saidat least one sensor (18), the error signal of VVD (E_(VVD))representative of a difference between a current value of the VVD and anoptimum value of VVD (VVD_(opt)); a closed-loop regulator (36) forcontrolling the AVD, receiving as input said error signal of AVD(E_(AVD)) and outputting a signal of AVD; and a closed-loop regulator(46) for controlling the VVD, receiving as input said error signal ofVVD (E_(VVD)) and outputting a signal of VVD.
 2. The device of claim 1,wherein said closed-loop regulators (36, 46) for controlling the AVD andthe VVD further are PID regulators.
 3. The device of claim 1, whereinone of said means for generating an error signal of AVD or VVD comprisesmeans for modulating (38, 48) and demodulating (42, 52) the AVD or theVVD, respectively.
 4. The device of claim 1, wherein: said means forgenerating an error signal of AVD comprises means for modulating (38)and demodulating (42) the AVD; said means for generating error signal ofVVD comprises means for modulating (48) and demodulating (52) the VVD;and said means for modulating (38, 48) and demodulating (42, 52) the VVDand of the AVD are functionally interdependent means.
 5. The device ofclaim 4, wherein either of said means for modulating (38, 48) anddemodulating (42, 52) the VVD and the AVD or the regulators (36, 46) forcontrolling the AVD and of the VVD further comprise means forcontrolling alternately the modulating and demodulating the AVD and theVVD over a predetermined number of cardiac cycles, and wherein theregulator for control of the AVD is inoperative during themodulation/demodulation of the VVD, and vice versa.